The ability to measure the three-dimensional structure of biological tissues is important, but common methods used to leave out basic information about the molecular composition and metabolic behavior of the tissue imaged. This molecular composition and metabolic behavior information may yield valuable scientific data on the behavior of biological systems, and would be of great clinical diagnostic value for finding diseases such as cancer. Much of the focus of biological and medical imaging today is to gain information about composition.
Functional magnetic resonance imaging (FMRI) utilizes contrast agents specific to particular molecular species or metabolic processes to provide specific information about the three-dimensional (3-D) location of these species or processes. This method is quite versatile, and poses relatively little risk to the patient, but is limited in practice to a resolution over 100 microns. Positron Emission Tomography (PET) and Single Photon Emission Computed Tomography (SPECT) utilizes a radiolabeled metabolic molecule to provide 3-D measurements of utilization of these molecules by sensing their gamma-ray emissions. Unfortunately, these methods have been cost prohibitive as diagnostic tools, and expose the patient to ionizing radiation. Fluorescence microscopy labels relevant biological structures and processes with a fluorescent compound which can be measured by conventional microscopy or confocal microscopy. The two-photon variant uses ultrafast pulses so that emission and excitation frequencies can be clearly separated, and also utilizes the additional selectivity that the intensity-sensitive nature of two-photon excitation provides. While some biological structures produce natural fluorescence, in general externally introduced fluorescence markers must be used, which can interfere with biological processes and are often toxic.
Molecules frequently have molecular resonance frequencies that are due to the electromagnetic attractions of atoms in the molecule. These frequencies are those of molecular vibrations, molecular rotational motions, the excitation of electrons to higher energy states, and occasionally finer structures such as hyperfine interactions and optical-magnetic properties. These properties are present without the introduction of any external contrast molecule. These frequencies are usually in the mid-infrared, corresponding to photons of 1.5-50 microns of wavelength. Unfortunately, they cannot be directly excited by electromagnetic radiation of the same frequency because when they are in tissue, the surrounding water absorbs almost all of these frequencies. The range of wavelengths that the tissue is relatively transparent to is 0.6-1.5 microns. Therefore multiphoton nonlinear processes need to be employed to probe these resonances. The photons to stimulate and record the processes are typically in a region where the tissue is not absorbing, so that they can reach the tissue feature and be measured from the feature.
Raman spectroscopy, first discovered in 1928, uses molecular resonance features of frequency Δω to split a photon of frequency ω into another photon of frequency ω−Δω and a resonance excitation of frequency Δω. The presence of photons at frequency ω−Δω identifies the concentration of the resonance feature. This process is in practice very weak and requires large amounts of power to produce any detectable amounts of photons. This weakness is due to the fact that the probability of a Raman excitation process to occur is proportional to the number of photons at frequency ω−Δω already present, of which there are typically few or none. Since photons that would be emitted by Raman excitation at frequency ω−Δω are indistinguishable from the incoming radiation that stimulates them, this is not a viable technique for achieving molecular sensitivity.
Coherent Anti-Stokes Raman Scattering (CARS) is another nonlinear spectroscopy technique that unlike conventional Raman spectroscopy, allows all of the photons necessary to stimulate the process to be introduced into the tissue by the illuminating source. This enables the probability of a CARS interaction to be increased to a (theoretically arbitrarily) high level so that a sufficient number of photons can be produced as to enable detection within a reasonable time period. It is essentially two stimulated Raman scattering processes in parallel. Two photons, a “pump” of frequency ω1 and a “Stokes” of frequency ω2 illuminate the tissue. They must be separated in frequency by ω1−ω2=Δω, which is the frequency of the molecular resonance. When molecules of the target molecular species are present, the resonance will be excited, and the pump photon will be converted to the same frequency as the Stokes photon. This is the first stimulated Raman scattering process. Another photon may arrive at frequency ω3 that will stimulate the emission of the excitation from the resonance, so that the energy of the photon of frequency ω4 and the excitation are converted to a new photon of frequency ω4=ω3+Δω, called the “anti-Stokes” photon. The presence of this photon of frequency ω4 indicates that a CARS process has taken place and indeed a molecule with the resonance feature is present. Often the “pump” beam is used as the photons of frequency ω3, so that ω3=ω1 and ω4=2ω1−ω2. Since the photon of ω4 is not the same frequency as one of the illuminating photons, and is typically within the transparency range of the tissue, it is easily discriminated from the incoming radiation. FIG. 1A shows an energy-level diagram for CARS, and FIG. 1B shows an energy-level diagram for Coherent Stokes Raman Scattering.
CARS microscopy uses the CARS process to look for the presence of a molecular species, but does not require any foreign substances to be introduced into the tissue. It scans the illumination point-by-point through the tissue and measures the number of generated anti-Stokes photons. When a three-dimensional mesh of points has been scanned, a complete three-dimensional picture of molecules of that resonance can be shown. Since CARS is a nonlinear process (and therefore is intensity sensitive), efficient conversion only occurs at the focus of the illumination, which can be made very tight (typically a half micron in both the axial and lateral directions). Therefore the resolution can be made many orders of magnitude better than MRI, which is the probably the largest competition for clinical use for similar purposes. Unfortunately, the penetration is usually rather low (limited to about 500 microns). A further shortcoming is that CARS microscopy measures the total number of anti-Stokes photons, or power, from the sample. However, the optical field contains temporal structure in the phase that is averaged out by power detection because photodetector response time is orders of magnitude slower than the oscillations of the optical field. The time scale on which the optical pulse varies (which is typically picoseconds or femtosecond time scales) is far too fast for photon detection equipment or electronics to detect (the fastest of which may detect 25 ps time scales).
Optical coherence tomography (OCT) is an emerging high-resolution medical and biological imaging technology [15-21]. OCT is analogous to ultrasound B-mode imaging except reflections of low-coherence light are detected rather than sound. OCT detects changes in the backscattered amplitude and phase of light.
Cross-sectional OCT imaging is performed by measuring the backscattered intensity of light from structures in tissue. This imaging technique is attractive for medical imaging because it permits the imaging of tissue microstructure in situ, yielding micron-scale imaging resolution without the need for excision and histological processing. Because OCT performs imaging using light, it has a one- to two-order-of-magnitude higher spatial resolution than ultrasound and does not require contact with tissue.
OCT was originally developed and demonstrated in ophthalmology for high-resolution tomographic imaging of the retina and anterior eye [22-24]. Because the eye is transparent and is easily optically accessible, it is well-suited for diagnostic OCT imaging. OCT is promising for the diagnosis of retinal disease because it can provide images of retinal pathology with 10 μm resolution, almost one order-of-magnitude higher than previously possible using ultrasound. Clinical studies have been performed to assess the application of OCT for a number of macular diseases [23,24]. OCT is especially promising for the diagnosis and monitoring of glaucoma and macular edema associated with diabetic retinopathy because it permits the quantitative measurement of changes in the retinal or retinal nerve fiber layer thickness. Because morphological changes often occur before the onset of physical symptoms, OCT can provide a powerful approach for the early detection of these diseases.
Recently, OCT has been applied for imaging a wide range of nontransparent tissues [16,17,25-27]. In tissues other than the eye, the imaging depth is limited by optical attenuation due to scattering and absorption. A “biological window” exists in tissue where absorption of near-infrared wavelengths is at a minimum and light can penetrate deep into highly-scattering tissue (FIG. 15) [28]. Because optical scattering decreases with increasing wavelength, OCT in nontransparent tissues has routinely used 1.3 μm wavelength light for imaging. In most tissues, imaging depths of 2-3 mm can be achieved using a system detection sensitivity of 110 dB (1 part in 1011). OCT has been applied to image arterial pathology in vitro and has been shown to differentiate plaque morphology with superior resolution to ultrasound [17,29].
Imaging studies have also been performed to investigate applications in gastroenterology, urology, and neurosurgery [30-32]. High resolution OCT using short coherence length, short-pulse light sources, has also been demonstrated and axial resolutions of less than 5 μm have been achieved [33,34]. High-speed OCT at image acquisition rates of 4 to 8 frames per second for 500 to 250 square pixel images has been achieved [35]. OCT has been extended to perform Doppler imaging of blood flow and birefringence imaging to investigate laser intervention [36-38]. Different imaging delivery systems including transverse imaging catheters and endoscopes, and forward imaging devices have been developed to enable internal body OCT imaging [39,40]. Most recently, OCT has been combined with catheter-endoscope-based delivery to perform in vivo imaging in animal models and human patients [41-44].
Apart from medical applications, OCT has been demonstrated as an emerging investigational tool for cell and developmental biology. OCT has imaged the development of numerous animal models including Rana pipiens and Xenopus laevis (Leopard and African frog), and Brachydanio rerio (zebrafish) [45-46]. High-speed OCT imaging has permitted the morphological and functional imaging of the developing Xenopus cardiovascular system, including changes in heart function following pharmacological interventions [47]. High-resolution imaging has permitted the real-time tracking of cell dynamics in living specimens including mesenchymal cell mitosis and neural crest cell migration [48]. OCT is advantageous in microscopy applications because repeated non-invasive imaging of the morphological and functional changes in genetically modified animals can be performed overtime without having to histologically process multiple specimens. The high-resolution, cellular-imaging capabilities suggest that OCT can be used to diagnose and monitor early neoplastic changes in humans.
The ability of OCT to perform optical biopsies, the in situ imaging of tissue microstructure at near-histological resolution, has been used to image morphological differences between normal and neoplastic tissue. OCT images of in vitro neoplasms of the female reproductive tract [49], the gastrointestinal tract [50], and the brain [51] have been investigated. Optical differences between normal and neoplastic tissue were evident, but primarily for late-stage changes. Still, situations exists were no inherent optical contrast exists between normal and pathologic tissue, such as in early-stage, pre-malignant tumors or in tumors which remain optically similar to normal tissue.
In the past, OCT has found numerous medical and biological applications. However, the imaging technique has relied largely on the inherent optical properties of the tissue to provide contrast and differentiate normal from pathological tissue. Phospholipid-coated perfluorobutane microbubbles (ImaRx Pharmaceutical, Tucson, Ariz.) have been used as a contrast agent for OCT; although they produce a strong OCT signal, blood and tissue also produce a fairly strong OCT signal, and the effects of this contrast agent in vivo on the visualization of blood vessels are subtle.